The present invention relates generally to implantable medical devices and more particularly to controlling surface properties of implantable biocompatible materials suitable for fabrication of implantable medical devices.
Implantable medical devices are fabricated of materials that are sub-optimal in terms of the biological response they elicit in vivo. Many conventional materials used to fabricate implantable devices, such as titanium, polytetrafluoroethylene, silicone, carbon fiber and polyester, are used because of their strength and physiologically inert characteristics. However, tissue integration onto these materials is typically slow and inadequate. Certain materials, such as silicone and polyester, elicit a significant inflammatory, foreign body response that drives fibrous encapsulation of the synthetic material. The fibrous encapsulation may have significant adverse effects on the implant. Moreover, conventional biomaterials have proved inadequate in eliciting a sufficient healing response necessary for complete device integration into the body. For example, in devices that contact blood, such as stents and vascular grafts, attempts to modify such devices to promote endothelial cell adhesion may have a concomitant effect of making the devices more thrombogenic.
When implanted, conventional blood-contacting implantable devices, such as stents, stent-grafts, grafts, valves, shunts and patches, fail to develop a complete endothelial layer, thereby exposing the device material to thrombus formation or smooth muscle cell proliferation, and ultimate failure of the implanted device. It has been recognized that, when implanted into the body, metals are generally considered to have superior biocompatibility than polymers used to fabricate commercially available polymeric grafts.
In investigating cellular interactions with prosthetic material surfaces, it has been found that cell adhesion to the material surface is mediated by integrins present on cell membranes that interact with the prosthetic surface. Integrins are the most prominent member of a class of extracellular matrix (ECM) adhesion receptors. Integrins are a large family of heterodimeric transmembrane proteins with different α and β subunits. Integrins are regulated at several levels. Modulation of the affinity of the adhesion receptor for ligand, termed affinity modulation, is a mechanism for activation of platelet aggregation and is believed to underlie activation of leukocyte adhesion. Adhesive strengthening by clustering of adhesion receptors or by cytoskeletal-dependent processes such as cell spreading has been shown to be crucial for strong cellular attachment, control of cell growth and cell motility. Under high shear forces present in flowing blood, leukocytes first tether, then roll along the vessel surface. When a local signal, e.g., a cytokine, is released in their vicinity, the leukocyte arrests, develops a firm adhesion then migrates across the endothelium. Tethering, rolling, arrest and adhesion tightening are all known to result from activation of leukocyte integrins.
Once adhered to a surface, cell spreading and migration are associated with assembly of focal adhesion junctions. Cell migration entails the coordination of cytoskeletal-mediated process extension, i.e., filopodia and lamellopodia, formation of adhesive contacts at the leading edge of a cell, breaking adhesive contacts, and cytoskeletal retraction at the trailing edge of the cell. Focal adhesions are comprised of integrins as the major adhesion receptors along with associated cytoplasmic plaque proteins. Assembly of focal adhesions is regulated by extracellular ligand binding events and by intracellular signaling events. Ligand binding controls localization of β1- and β3-containing integrins into focal adhesions. The cytoplasmic domains of the β subunits have intrinsic signals for focal adhesion localization, but incorporation of the integrins into focal adhesions is prevented by the α subunits of the heterodimers. Ligand binding, however, relieves this inhibition and allows the subunit cytoplasmic tail signals to recruit the integrin dimmer into the focal adhesion.
Attempts at coating implanted metal devices, such as stents, with proteins that contain the Arg-Gly-Asp (RGD) attachment site have been made with some success. The RGD sequence is the cell attachment site of a large number of adhesive extracellular matrix, blood, and cell surface proteins and many of the known integrins recognize the RGD sequence in their adhesion protein ligands. Integrin-binding activity may also be reproduced by synthetic peptides containing the RGD sequence. However, bare metal implanted materials will not, of course, have native RGD attachment sites. Thus, metal implantable devices, such as stents, have been derivitized with polymers having RGD attachment sites bound to the polymer matrix.
It has been found that when prosthetic materials are implanted, integrin receptors on cell surfaces interact with the prosthetic surface. When cells come into contact with the extracellular matrix, such as a prosthetic surface, their usual response is to extend filopodia, and integrins at the tip of the filopodia bind to the extracellular matrix and initiate the formation of focal adhesions. Actin-rich lamellipodia are generated, often between filopodia, as the cell spreads on the extracellular matrix. Fully developed focal adhesions and associated actin stress fibers ensue. These same evens occur during cell migration as cells extend lamellipodia and form focal adhesions to derive the traction necessary for movement. Giancotti, F. G., et al. Science, 285:13 August 1999, 1028-1032.
The integrin receptors are specific for certain ligands in vivo. If a specific protein is adsorbed on a prosthetic surface and the ligand exposed, cellular binding to the prosthetic surface may occur by integrin-ligand docking. It has also been observed that proteins bind to metals in a more permanent fashion than they do to polymers, thereby providing a more stable adhesive surface. The conformation of proteins coupled to surfaces of most medical metals and alloys appears to expose greater numbers of ligands and attract endothelial cells having surface integrin clusters to the metal or alloy surface, preferentially over leukocytes.
Because of their greater adhesive surface profiles, metals are also susceptible to short-term platelet activity and/or thrombogenicity. These deleterious properties may be offset by administration of pharmacologically active antithrombogenic agents in routine use today. Surface thrombogenicity usually disappears 1-3 weeks after initial exposure. Antithrombotic coverage is routinely provided during this period of time for coronary stenting. In non-vascular applications such as musculoskeletal and dental, metals have also greater tissue compatibility than polymers because of similar molecular considerations. The best article to demonstrate the fact that all polymers are inferior to metals is van der Giessen, W J. et al. Marked inflammatory sequelae to implantation of biodegradable and non-biodegradable polymers in porcine coronary arteries, Circulation, 1996:94(7):690-7.
Normally, endothelial cells (EC) migrate and proliferate to cover denuded areas until confluence is achieved. Migration, quantitatively more important than proliferation, is affected by exposure of the EC to blood flow. Under static conditions or in the presence of minor shear stress, for example, about 1.5 dynes/cm2, EC have been observed to migrate at speeds between about 10 μm/hr to about 15 μm/hr. Palmaz, J. C., Bailey, S., Marton, D., and Sprague, E. Influence of stent design and material composition on procedure outcome J. Vasc. Surg. 2002; 36:1031-1039. Also, the cause of restenosis includes vessel injury due to pressure from stent expansion and neointimal thickening due to decrease in vessel wall shear stress (WSS).
EC migrate by a rolling motion of the cell membrane, coordinated by a complex system of intracellular filaments attached to clusters of cell membrane integrin receptors, specifically focal contact points. The integrins within the focal contact sites are expressed according to complex signaling mechanisms and eventually couple to specific amino acid sequences in substrate adhesion molecules. An EC has roughly 16-22% of its cell surface represented by integrin clusters. Davies, P. F., Robotewskyi A., Griem M. L. Endothelial cell adhesion in real time. J. Clin. Invest. 1993; 91:2640-2652, Davies, P. F., Robotewski, A., Griem, M. L., Qualitiative studies of endothelial cell adhesion, J. Clin. Invest. 1994; 93:2031-2038. This is a dynamic process, which involves more than 50% remodeling in 30 minutes.
The focal adhesion contacts vary in size and distribution, but 80% of them measure less than 6 μm2, with the majority of them being about 1 μm2, and tend to elongate in the direction of flow and concentrate at leading edges of the cell. Although the process of recognition and signaling to determine specific attachment receptor response to attachment sites is not completely understood, availability of attachment sites will favorably influence attachment and migration. It is known that materials commonly used as medical grafts, such as polymers, do not become covered with EC and therefore do not heal after they are placed in the arteries. Furthermore, heterogeneities of materials in contact with blood flow are preferably controlled by using vacuum deposited materials.
There have been numerous attempts to increase endothelialization of implanted medical devices such as stents, including covering the stent with a polymeric material (U.S. Pat. No. 5,897,911), imparting a diamond-like carbon coating onto the stent (U.S. Pat. No. 5,725,573), covalently binding hydrophobic moieties to a heparin molecule (U.S. Pat. No. 5,955,588), coating a stent with a layer of blue to black zirconium oxide or zirconium nitride (U.S. Pat. No. 5,649,951), coating a stent with a layer of turbostratic carbon (U.S. Pat. No. 5,387,247), coating the tissue-contacting surface of a stent with a thin layer of a Group VB metal (U.S. Pat. No. 5,607,463), imparting a porous coating of titanium or of a titanium alloy, such as Ti—Nb—Zr alloy, onto the surface of a stent (U.S. Pat. No. 5,690,670), coating the stent, under ultrasonic conditions, with a synthetic or biological, active or inactive agent, such as heparin, endothelium derived growth factor, vascular growth factors, silicone, polyurethane, or polytetrafluoroethylene (U.S. Pat. No. 5,891,507), coating a stent with a silane compound with vinyl functionality, then forming a graft polymer by polymerization with the vinyl groups of the silane compound (U.S. Pat. No. 5,782,908), grafting monomers, oligomers or polymers onto the surface of a stent using infrared radiation, microwave radiation or high voltage polymerization to impart the property of the monomer, oligomer or polymer to the stent (U.S. Pat. No. 5,932,299). However, all these approaches do not address the lack of endothelialization of polymer grafts.
Overall rate to reach confluence for the endothelial cells on the blood contact surface of implanted medical device is mainly determined by two factors, the rate of cell movement and rate of cell proliferation, with the first being more important. The rate of cell movement further comprises three interrelated steps. Initially, a cell forms lamellipodia and filopodia that protrude outward. This step involves reassembly of actins in the forefront of lambaepolia. After protrusion of lamellipodia from one or multiple points from the cell membrane, the front end of the lamellipodia will form a close attachment, called focal adhesion point, to the substratum through the interaction of integrin on the cell membrane and extracellular matrix binding site. The final step of cell movement involves the contraction of the posterior end through the action of myosin II. The formation of a focal adhesion point is critical for the cell movement because the protruding lamellipodia will otherwise fold back. Without the tension force from the focal adhesion point, a cell loses the contraction from the posterior end and hence stops moving.
Availability of attachment sites on the substratum is not only important for the focal adhesion point formation, but also important for propagation. It has been shown that when cells are forced to spread, they survive better and proliferate faster than cells that are confined to the same amount of surface area (Science 276:1425-1428, 1997). This may explain why spreading of neighbor cells stimulate a cell to proliferate, after cells are lost from epithelium.
The formation of extracellular matrix (ECM) is, to much extent, determined by the cells within it because molecules which form ECM are secreted by the cells. Subsequently, the structure of the ECM, and hence the distribution of attachment sites on the ECM for the integrin binding, determines the focal adhesion point formation, the critical step in cell movement. Therefore, proper distribution of integrin binding sites on the surface of an implanted medical device substantially determines the speed of reendothelialization from the ends surrounding the device.
There still remains a need for a medical device that stimulates endothelial proliferation and movement when implanted in order to form an endothelial layer over the medical device. Furthermore, there is a remaining need for a method of fabricating such a medical device.